Digital X-ray sensors exist comprising a conversion layer in the form of an amorphous coating, normally made of Amorphous Selenium or of Cesium iodide, and an integration panel, i.e. a collection layer, that has a TFT pixel structure (Thin Film Transistor). The conversion layer serves for transforming into an electric charge the photons of an X-ray beam that has travelled across an irradiated sample. This may occur directly or indirectly, as in the case of amorphous Selenium and of Cesium iodide; respectively. The total charge obtained by the conversion during an X-ray exposure builds up in the pixels of the integration panel.
Once the exposure has been completed, the amount of charge accumulated in each pixel is read. More in detail, an image acquisition electronics is provided that comprises an analog-to-digital converter arranged at the boundary of the integration panel. The analog-to-digital converter changes the overall charge accumulated in each pixel into an electric voltage, i.e. into a number that is proportional to the overall radiation that has travelled across the sample at each pixel of the integration panel. These numbers can be converted into a radiographic image in which the contrast depends upon the overall radiation that is accumulated in each pixel.
A so-called “photon-counting” technique is also known, in which the photons are counted one-by-one, and are ranked into a plurality of channels, thus obtaining a “film grade” resolution, i.e. a resolution that is comparable with the resolution allowed by high-resolution radiographic plates. In particular, hybrid detectors exist that are known as MediPix and that are provided with an ASIC for carrying out a photon-counting procedure. These hybrid detectors comprise discriminators associated with event counters that are used in such a way that the image acquisition electronics counts only events, i.e. acquisitions of photons that fall in a predetermined energy window. This way, an X-ray imaging technique is obtained that has spectroscopic features. A more recent device, known as Medipix-3, has a finer energy resolution thanks to a real-time charge share correction. Medipix 3 also comprises multiple pixel counters that can be used in different operation modes. This allows a continuous detection, and up to eight energy thresholds can be obtained.
In the Medipix device, like in other devices, the collection layer is implemented by CMOS technology, which is a low-power consumption technology, i.e. about a few Watts, and a low-cost technology. For this reason, the CMOS technology is preferred, in particular, to provide ASICs consisting of a large number of pixels, e.g. about 106, as it is required for a radiological device.
However, as well known, CMOS-based devices are inherently non-calibrated devices. In other words, the units consisting of discriminators/counters of this type cannot provide homogeneous counting responses. In fact, each CMOS counting chain has its own offset value that, moreover, may depend upon environmental conditions such as the temperature, and other operating conditions. A counter associated with a discriminator that has an offset value nearer to the threshold associated with it counts more events than a counter that is associated with a discriminator that has an offset farther from the threshold of the same nominal value. For this reason, an ASIC suitable for the invention, which is implemented by CMOS technology, can give rise to a non-homogeneous response, within a same pixel and from a pixel to another.
Another remarkable drawback of the prior art is that various sources of noise exist. In addition to quantum noise, which is unavoidably related to the intensity of the X-ray beam, a certain level of additional noise is always present, which depends upon the sensor and upon the detection electronics, in particular it depends upon the means used for amplifying the charge. Total, quantum and additional noise prevents the collection pixels from detecting the charge delivered by the conversion layer, when the charge is lower than a predetermined threshold. Therefore, the need is felt that each pixel should contain as much charge as possible, and/or that the noise should be substantially reduced to the quantum level only, so that each pixel can work if it is reached by a minimum charge amount.
In the light of the above, a digital sensor formed by pixels must carry out corrections, or calibrations, in order to limit or avoid some drawbacks due to the sensor itself or to the technology on which the read electronics (ASIC) is based. The main effects that must be corrected are:                the occurrence of a “dark current” generated by the sensor, i.e. an intrinsic current that can be detected in the absence of any outer irradiation and of strong collector electric fields. The dark current is one of the most important components of additional noise, in addition to quantum noise;        an offset of the direct current level at the outlet of the inlet stage, i.e. of the pre-amplification stage, and of the gain stage of the pixel electronics;        an offset of the direct current level at the inlet of the discriminator/counter unit.        
Sindre Mikkelsen et al.1 describe a radiography sensor comprising an X-ray-sensitive conversion layer and a 64-pixel ASIC collector, provided with a photon-counting function. Each pixel of the ASIC has a single inlet terminal that is normally connected to an electrode of the radiation sensor. In an example, which discloses a calibration function, each pixel is connected to a common calibration node through a switch controlled by a software outside of the ASIC, and a calibration network is provided, i.e. an adjustment means, that enables a user to switch the inlet terminal of each pixel from a pad of the sensor pixel to the calibration node. Therefore, this sensor is configured for executing a sequential calibration, i.e. a pixel-by-pixel calibration, under the control of the software. Such a calibration technique cannot practically be used in the case of a matrix that has a large number of pixels, since the pixel-by-pixel calibration would need a too long calibration time, in comparison to the requirements, for instance, of a radiological sensor that must carry out a large number of radiological sessions within a prefixed time.
Dinapoli et al.2, and Radicci et al.3 describe a similar radiography sensor provided with a manual calibration circuit. This circuit is configured for working in a test mode, in which a known amount of charge is supplied to the inlet of an amplifier of a predetermined pixel. The response of the amplifier of the selected pixel is available to be shown by a display device like an oscilloscope. In particular, a precise calibration of the chip is carried out with a monochromatic X-ray source once the collection chip has been connected to the photon-sensitive layer by a bump-bonding technique. Also in this case, the calibration is carried out pixel-by-pixel. Moreover, the calibration has to be carried out manually by an operator, and, consequently, the technique cannot be readily used for matrices that comprise a large number of pixels.
Perenzoni et al.4 describe a reading circuit configured for carrying out a photon-counting function for a radiography sensor formed by pixel, in which a completely analog self-calibration procedure is provided. An analog self-calibration must be repeated before each photon capture event in the same conditions as in the normal operation. Therefore, the noise that occurs during the calibration procedure has the same level as the noise that occurs while receiving the data from the sample, so the noise of the calibration procedure is added to the noise of the normal operation.
Devices like the above-mentioned Medipix have further drawbacks. In particular, they do not allow making a surface larger than 14×14 mm, whereby 256×256 pixels can be provided at most. Furthermore, the pixels are arranged in a square grid. Such an arrangement is not the best suited for sampling the radiation field, since it leads to a worse sampling out of the directions of a couple of orthogonal axis. Finally, Medipix has been conceived to be preferably coupled with Silicon converters, in which the photon-to-charge conversion is carried out at a very low efficiency if the photon energy is higher than 15 keV. These drawbacks do not allow using such devices as Medipix to make medical radiology sensors. Briefly, sensors cannot be provided that have a continuous active surface large enough and a sensitivity to X-ray energy high enough for this purpose.
Another important drawback of the prior art is the low conversion efficiency of the conversion layer. Selenium converters are known that have a thickness not larger than 0.5 mm. These converters, due to the relatively low atomic number of Selenium, are sensitive to X-ray beams up to a maximum energy of about 30-40 keV only. Further Caesium iodide converters, even if manufactured with a suitable thickness, are however affected by a resolution loss, due to the indirect X-ray beam-to-charge conversion process. It is therefore desirable to provide a digital X-ray sensor that has a high efficiency, and a low-noise conversion layer, by which a “film grade” resolution can be obtained.
Further, so-called integration techniques are available on surfaces that allow their use in medical radiology, as in the case of WO96/33424. However, such integration techniques are not suitable for ranking the photons converted during a same flash according to the energy that the photons possess when they reach the conversion layer. This feature, which is also known as the “colour” of the photons, can be very relevant in some diagnostic and analytic procedures. By the known integration techniques, however, the “colour” can be obtained by carrying out so many flashes as the number of the energy levels of interest, and by changing in turn the energy of the X-ray beam. Besides taking a long time and requiring a large amount of resources, this technique would expose a living subject to a strong dose of harmful radiation.